The field of the invention relates to systems and methods for magnetic resonance imaging (“MRI”). More particularly, the present invention relates to systems and methods controlling a radio frequency (RF) circuit for use with an MRI system.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0) applied along, for example, a z axis of a Cartesian coordinate system, the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but process about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A NMR signal is emitted by the excited spins after the excitation signal B1 is terminated, this signal may be received and processed to form an image or produce a spectrum.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
Radio frequency antennas, or coils are used to produce the excitation field B1 and other RF magnetic fields in the subject being examined. Such coils are also used to receive the relatively weak NMR signals that are produced in the subject. Such coils may be so-called “whole body” coils that are large enough to produce a magnetic field for a human subject or, they can be much smaller “local” coils that are designed for specific clinical applications such as head imaging, knee imaging, wrist imaging, and the like. Local coils may be either volume coils or surface coils.
The aforementioned polarizing magnetic field is a common metric upon which standard systems are differentiated. Standard magnetic field strengths include 1.5 Tesla (T), 3 T, as well as those of lesser and greater strength. Increased magnetic field strength brings better signal-to-noise ratio (SNR), higher resolution, and improved contrast and, therefore, experimental system use ultra-high-fields of 7 T, 9.4 T and 11.74 T.
MR Imaging at higher magnetic fields strengths, including the above-referenced ultra-high-fields, presents certain challenges in RF coil circuit design. The common RF transverse electromagnetic (TEM) coil design has widely used microstrip transmission line as elements that inductively couple to the human anatomy at Larmor frequencies of up to 500 MHz (11.74 T). As used herein RF coil, RF antenna, microstrip, and the like all refer generally to electrical elements or “RF elements” and are used herein. As shown in FIG. 1A, the general RF structure for an MRI system is illustrated and includes an RF coil element 10 on a dielectric substrate 12, for example Teflon, wireless receiver components 14, wireless transmitter components 16, an RF switch 18, and capacitors 20.
RF coils 10 for use with a microstrip line provide advantages, including distributed coil circuit, high sensitivity due to high Q, and relatively-simple structure. However, this high sensitivity also creates a critical disadvantage in the form of a loading (body) effect. As illustrated in FIG. 1B, the resonance frequency and quality factor (Q) are changed from location 22 to location 24 due to impedance mismatch when different human body weight, shape, and tissue composition are loaded. As illustrated, resonance frequency shifts down from the Larmor frequency determined by the strength of magnetic field (B0) because of the coupling between RF coil and human anatomy.
The loading effect needs to be taken into consideration by way of a tuning procedure after the body comes into the MRI scanner and it is unpredictable. In general, referring again to FIG. 1A, tuning of the RF coil system entails adjusting the capacitors 20. A first capacitor is called the matching capacitor (Cm) connected in series and another capacitor called the tuning capacitor (Ct) connected in parallel. The matching capacitor Cm matches the impedance of the RF coil together with the effects of human anatomy to the source and power amplifiers. The tuning capacitor Ct holds the resonance frequency (which is the Larmor frequency) of the RF coil element, which is determined by the magnetic field strength (B0) by:
                                          ω            0                    =                      γ            ×                          B              0                                      ,                                  ⁢                              (                          γ              =                                                42.58                  ⁢                                                                          ⁢                  MHz                                Tesla                                      )                    .                                    Eqn        .                                  ⁢                  (          1          )                    
In current practice essentially all coils operating in transmit, receive, and transceiver modes in MRI applications operate in fixed tuned and fixed impedance matched conditions. The isolation between coil elements, between transmit and receive coils, or modes is also fixed. Coils are designed, manufactured, and used to be “one size fits all”. However all human body loads to which coils are applied are not the same size, are not the same shape, are not in the same position and, therefore, do not present the same electrical load to the fixed coils. Because these coils with fixed tune, match, and isolation conditions cannot be adjusted by any existing practical means, suboptimal coil performance is the consequence. Reflected power lost to impedance mismatch, attenuated power lost to off-resonance transmission and reception, field distortions, and power loss to coil-to-coil and T/R mode coupling (lack of isolation) renders images with lower signal-to noise ratios, lower homogeneity, more RF artifacts, and higher specific absorption rates.
These problems have been tolerated at lower field strengths, such as 1.5 T and below, because the longer wavelengths for the Lower Larmor frequencies produce fields with stronger penetration and higher uniformity, attributes which compensate somewhat for the problems of ignoring coil tuning, matching and isolation. However at higher B0 fields and B1 frequencies, coil tuning and matching becomes more critical for the reasons given above. And while, at lower frequencies, single monolithic resonators can be used to generate uniform excitation fields, safe and successful images at higher fields increasingly benefit from multi-channel transmit, receive, and transceiver coils. Multichannel coils give the ability to adjust the B1 field in any or all of the phase, magnitude, frequency, space, or time domains to facilitate B1 field optimization over a field of interest. Each channel of a multi-channel coil must be tuned, matched, and isolated. Also, each channel should, ideally, be tuned, matched, and isolated per patient or other load. That is, it would be beneficial if each channel were tuned, matched, and/or isolated for each patient, or, even better, be tuned, matched and/or isolated dynamically to track patients' movements over the imaging process, be it course physical movements of the body, or be it breathing, heartbeat, or other physiological motion. Given that receivers of up to 64 channels and transmitters of up to 16 channels are being delivered with MRI systems today, manual adjustment of tune, match, and/or isolation capacitances per channel is impractical for either clinical or research applications. That is, practically speaking, to make such adjustments, the operator of the MRI scanner would need to adjust the capacitances of these capacitors 20 by hand. It is a major obstacle to the application of these coils to the MRI system.
Therefore, it would be desirable to have a system and method for providing and operating an RF system within an MR imaging process that does not require cumbersome tuning, matching, and adjustments thereto that varying substantially with operational characteristics of the MR system and the subject being imaged.